1. Field of the Invention
The present invention generally relates to a magnetic resonance imaging system employing active shield gradient coils for magnetically canceling a leakage gradient field. More specifically, the present invention is directed to an active shield gradient coil system to effectively mitigating eddy current loss occurring in heat-shield members for maintaining a superconducting coil assembly at very low temperatures.
2. Description of the Related Art
In FIG. 1, there is shown one of the conventional magnetic resonance (hereinafter referred to as "MR") imaging apparatuses. This conventional MR imaging apparatus comprises a gantry 101 having an imaging hole 101A at a center thereof. In this gantry 101, there are provided a superconducting magnet 102 around the peripheral portion thereof for producing a uniform static magnetic field and a gradient field coil 103 for producing a gradient magnetic field to be superimposed with the static magnetic field within the magnet bore 101A. The superconducting magnet 102 is so constructed that both a toroidal superconducting coil 104 stored within a liquid helium bath (not shown in detail), and a heat shield tube 105 for thermally shielding the superconducting coil 104 are stored in an outer vessel 106, the cross-sectional shape thereof being toroidal. A refrigerator 107 is employed so as to cool the superconducting magnet 5 by utilizing a vapored helium gas, whereby a total amount of vapored helium within the helium bath can be suppressed.
During MR imaging operation, an object under medical examination, e.g., a patient (not shown) is inserted into this magnet bore 101A, the static magnetic field produced from the superconducting magnet 102 is uniformly applied particularly to a specific portion (a portion to be imaged) of the patient, and also RF magnetic fields are applied thereto in a direction perpendicular to the application direction of the static magnetic field. Furthermore, since the preselected gradient magnetic fields produced from the above-described gradient field coil 103 are superimposed on the static magnetic field, the MR phenomenon may occur only at the specific slice portion of the object under medical examination and an MR signal (e.g., FID signal and spin echo signal) generated from the specific slice portion is acquired after the application of the RF fields has been accomplished. The acquired MR signal is further processed in the image data processor (not shown) by way of, e.g., Fourier transform, whereby tomographic images may be reconstructed and desirable tomographic images may be displayed on a monitor (not shown).
When the gradient magnetic fields are produced from the gradient magnetic coil 103, pulsatory leakage fields may happen to occur therefrom so that eddy currents will be produced on the heat shield tube 105 for maintaining the superconducting magnet 104 at a very low temperature, and furthermore, another pulsatory magnetic field is newly generated by these eddy currents, which will in turn be superimposed with the existing gradient fields. As a result, the resultant gradient magnetic fields cannot constitute a predesigned gradient magnetic field, i.e., magnetic flux and distribution shapes. Accordingly, artifact may be induced in the tomographic images and the signal level of the MR signal is lowered, resulting in deterioration of image quality.
To solve the above-described problems caused by the leakage field from the gradient field, a so-called "active shield method" has been proposed.
For instance, see U.S Pat. No. 4,737,716 issued to Roemer et al. on Apr. 12, 1988; U.S. Pat. No. 4,733,189 issued to Punchard et al. on Mar. 22, 1988; and U.S. Pat. No. 4,794,338 on Dec. 27, 1988.
A basic idea of such a conventional active shield method will now be summarized with reference to an illustration of FIG. 2. It should be noted that the same reference numerals employed in FIG. 1 are employed to denote the same or similar components shown in FIG. 2.
As shown in FIG. 2, an active shield gradient coil 110 is employed between the active shield gradient coil 107 and superconducting magnet 102 in order to cancel the above-described leakage field by superimposing a gradient shield field generated by this active shield gradient coil 110 onto the leakage field. A field direction of this gradient shield field is opposite to that of the leakage field. Accordingly, employment of the active shield gradient coil 110 may prevent the eddy current from being produced in the heat shield tube 105. In general, both the gradient field coil 107 and active shield gradient coil 110 are mounted within a bore (a typical diameter is designed for approximately 1,000 mm) formed at a center portion of the outer vessel 106, and both of these coils 107 and 110 are integrally fabricated.
However, there is another problem in the above-described active shield method. That is, since an interval between the gradient field coil 107 and active shield gradient coil 110 is rather short, or the gradient field coil 107 is positioned in close proximity to the active shield gradient coil 110, large energizing current must be supplied to the active shield gradient coil 110 so as to sufficiently cancel a gradient leakage field leaked from the gradient magnetic field of the active shield gradient coil 107. Therefore, total power consumption is increased and large heat dissipation will occur from the active shield gradient coil 110.
In addition, to reduce such large heat from the active shield gradient coil 110, another cooling means different from the above-described refrigerator 107 is additionally required.
As previously described, since two coil energizing currents having mutually opposite flow directions must be supplied to both coils 107 and 110 and thus a difference between the field strengths of these coils 107 and 110 becomes a strength of the gradient field actually applied to the patient (not shown), the following problem may occur. That is, the shorter an interval between two coils 107 and 110, the smaller a difference between the field strengths produced by these coils become, so that the strength of the actually applied gradient field becomes small in inverse proportion to current valves of these coils 107 and 110 and the desirable gradient field cannot be effectively produced.
To separate these coils 107 and 110 from each other as long as possible within this bore, the diameter of the bore may be enlarged. However, this causes such a drawback that the entire dimension of the superconducting magnet 102 becomes large, and therefore a large install space as well as higher cost are necessarily required. In a particular case, the existing examination room for the MR imaging purpose cannot be used and a new examination room with a higher ceiling height must be prepared for installing such large-sized MR imaging apparatus. Alternatively, the diameter of the gradient magnetic coil 107 may be made small. However, this cause another difficulty that a patient must be inserted into such a narrower space, which will in turn induce claustrophobia. Accordingly, the above-described solutions do not have yet been realized in the practical MR imaging field.
Furthermore, if the above-described active shield gradient coil is tried to be entered into the bore of the first-described conventional MR imaging apparatus shown in FIG. 1, the present diameter of this bore must be expanded and therefore huge modification cost and cumbersome workloads will be necessarily required.
In accordance with another practical solution, slits may be formed in the heat shield tube 105 so as to suppress the eddy current loss. However, the leakage field generated by the gradient field coil 107 may reach the above-described liquid helium tub positioned inside the heat shield tube 105 and then causes another eddy current loss in this liquid helium tube. As a consequence, a total amount of vapored liquid helium is increased due to this eddy current heat dissipation.